Magnetic resonance imaging with resolution and contrast enhancement

ABSTRACT

MRI scans typically have higher resolution within a slice than between slices. To improve the resolution, two MRI scans are taken in different, preferably orthogonal, directions. The scans are registered by maximizing a correlation between their gradients and then fused to form a high-resolution image. Multiple receiving coils can be used. When the images are multispectral, the number of spectral bands is reduced by transformation of the spectral bands in order of image contrast and using the transformed spectral bands with the highest contrast.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention is directed to a system and method for improvingthe resolution and tissue contrast in MRI.

2. Description of Related Art

The best current source of raw image data for observation of a complexsoft tissue and bone structure is magnetic resonance imaging (MRI). MRIinvolves the transmission of RF signals of predetermined frequency(e.g., approximately 15 MHZ in some machines, the frequency dependingupon the magnitude of magnetic fields employed and the magnetogyricratio of the atoms to be imaged). Typically, exciting pulses of RFenergy of a specific frequency are transmitted via an RF coil structureinto an object to be imaged. A short time later, radio-frequency NMRresponses are received via the same or a similar RF coil structure.Imaging information is derived from such RF responses.

In MRI, a common imaging technique is the formation of images ofselected planes, or slices, of the subject being imaged. Typically thesubject is located in the static magnetic field with the physical regionof the slice at the geometric center of the gradient field. Generally,each gradient will exhibit an increasing field strength on one side ofthe field center, and a decreasing field strength on the other side,both variations progressing in the direction of the particular gradient.The field strength at the field center will thus correspond to a nominalLarmor frequency for the MRI system, usually equal to that of the staticmagnetic field. The specific component of a gradient which causes thedesired slice to be excited is called the slice selection gradient.Multiple slices are taken by adjusting the slice selection gradient.

However, MRI often introduces the following technical challenges. Manyof the anatomical structures to be visualized require high resolutionand present low contrast, since, for example, many of themusculoskeletal structures to be imaged are small and intricate. MRIinvolves the use of local field coils to generate an electromagneticfield; such local field coils form a non-uniform illumination field. MRIimages can also be noisy.

In particular, MRI has the following limitations in resolution andtissue contrast. Although current MRI machines can achieve relativelyhigh intra-plane resolution, the inter-slice resolution is not so goodas the intra-plane resolution; also, the inter-slice resolution islimited by the ability of the system to stimulate a single spatial sliceor section. Although tissue contrast can be adjusted by selecting theright pulse sequence, analysis of a single pulse sequence is not enoughto differentiate among adjacent similar tissues. In other words, theresolution is typically poor in the out-of-plane dimension, and thecontrast is typically low between soft tissue structures.

SUMMARY OF THE INVENTION

It will be readily apparent from the foregoing that a need exists in theart to overcome the above-noted limitations of conventional MRI.

Therefore, it is an object of the present invention to increaseinter-slice resolution.

It is another object of the present invention to improve tissuecontrast.

It is still another object of the present invention to improveinter-slice resolution and tissue contrast simultaneously.

It is yet another object to provide a simple technique for imageregistration.

To achieve the above and other objects, the present invention isdirected to a system and method for creating high-resolution MRI volumesand also high-resolution, multi-spectral MRI volumes. At least oneadditional scan is obtained in an orthogonal direction. Then, through adata fusion technique, the information from an original scan and anorthogonal scan are combined, so as to produce a high-resolution, 3Dvolume. In addition, one may use a different pulse sequence in theoriginal orientation or in an orthogonal orientation, and a data fusiontechnique can be applied to register the information and then visualizea high-resolution, multi-spectral volume.

BRIEF DESCRIPTION OF THE DRAWINGS

A preferred embodiment will be set forth in detail with reference to thedrawings, in which:

FIG. 1 shows a block diagram of an MRI system according to the preferredembodiment;

FIGS. 2 and 3 show flow charts showing steps performed in registeringand fusing two images;

FIG. 4 shows a blurring effect caused by the correlation of the twoimages;

FIGS. 5A-5C show a typical signal, its gradient and a comparison of theautocorrelations of the signal and its gradient, respectively;

FIGS. 6A and 6B show two voxels scanned in orthogonal directions;

FIG. 6C shows the problem of deriving high-resolution information fromthe voxels of FIGS. 6A and 6B;

FIGS. 7A-7I show comparisons between the individual scans and the fusedimage;

FIGS. 8A-8I show a comparison among simple fusion without registration,simple fusion after registration and complete fusion;

FIGS. 9A-9I show fusion of orthogonal images without correlation;

FIGS. 10A and 10B show images taken with three and four local receivercoils, respectively;

FIGS. 11A and 11B show a two-band spectral image of a knee; and

FIGS. 11C and 11D show principal components of the image of FIGS. 11Aand 11B.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT

A preferred embodiment of the present invention will now be set forth indetail with reference to the drawings.

FIG. 1 shows a block diagram of an MRI system 100 on which the presentinvention can be implemented. The system 100 uses an RF coil 102 and agradient coil 104 to apply the required RF and gradient fields to thesubject S. A spectrometer 106, acting under the control of a computer108, generates gradient signals which are amplified by an X amplifier110, a Y amplifier 112 and a Z amplifier 114 and applied to the gradientcoil 104 to produce the gradient fields. The spectrometer 106 alsogenerates RF signals which are amplified by an RF amplifier 116 andapplied to the RF coil 102 to produce the RF fields. The free inductiondecay radiation from the sample S is detected by the RF coil 102 or byone or more local receiving coils 118 and applied to the spectrometer106, where it is converted into a signal which the computer 108 cananalyze.

The computer 108 should be sufficiently powerful to run a mathematicalanalysis package such as AVS, a product of Advanced VisualizationSystems of Waltham, Mass., U.S.A. Examples are the Apple Power Macintoshand any IBM-compatible microcomputer capable of running Windows 95, 98or NT. The significance of the local receiving coils 118, andparticularly of the number used, will be explained in detail below. Theother components of the system 100 will be familiar to those skilled inthe art and will therefore not be described in detail here.

The various techniques to enhance the images will now be described indetail.

Inter-Slice Resolution

The inter-slice resolution problem is solved by using two volumetricdata sets where scanning directions are orthogonal to each other, andfusing them in a single high-resolution image. FIG. 2 shows an overviewof the process. First, in step 202, a first scan of the subject istaken. In step 204, a second scan of the subject is taken in a directionorthogonal to that of the first scan. Third, in step 206, the images areregistered. Finally, in step 208, the images are fused. FIG. 3 shows thesteps involved in image registration. In step 302, gradients of theimage data are formed. In step 304, the gradients are correlated. Instep 306, the correlation is maximized through a hill-climbingtechnique.

Although the fusion of the two volumes seems to be trivial, two issueshave to be taken into account: 1) the registration among volumes scannedat different time (step 206) and 2) the overlapping of sampling voxelvolumes (step 208). Those issues are handled in ways which will now bedescribed.

Image Registration (Step 206)

The goal of image registration is to create a high-resolution 3D imagefrom the fusion of the two data sets. Therefore, the registration ofboth volumes has to be as accurate as the in-plane resolution. Thepreferred embodiment provides a very simple technique to register twovery similar orthogonal MRI images. The registration is done by assuminga simple translation model and neglecting the rotation among the twovolumes, thus, providing a fair model for small and involuntary humanmotion between scans.

An unsupervised registration algorithm finds the point (x, y, z) wherethe correlation between the two data sets is maximum. The Schwartzinequality identifies that point as the point where they match. Giventwo functions u(x, y, z) and v(x, y, z), the correlation is given by:r(x, y, z) = u(x, y, z) * v(x, y, z) = ∫∫∫u(α, β, γ)v(x + α, y + β, z + γ)𝕕α𝕕β𝕕γ.If u(x, y, z) is just a displaced version of v(x, y, z), or in otherwords, u(x, y, z)=v(x+Δx, y+Δy, z+Δz), then the maximum is at (−Δx, −Δy,−Δz), the displacement between the functions.

In MRI, every voxel in the magnetic resonance image observes the averagemagnetization of an ensemble of protons in a small volume. The voxel cantherefore be modeled as the correlation of the continuous image by asmall 3D window. Let g₁(x, y, z) and g₂(x, y, z) be the sampled volumesat two orthogonal directions, which are given by:g ₁(x,y,z)=f(x,y,z)*w ₁(x,y,z)*Π(x,y,z)g ₂(x,y,z)=f(x+Δx,y+Δy,z+Δz)*w ₂(x,y,z)*Π(x,y,z)where w₁(x, y, z) and w₂(x, y, z) are the 3D windows for the twoorthogonal scanning directions, (Δx, Δy, Δz) is a small displacement,and Π(x, y, z) is the sampling function. Therefore, the correlation ofthe two sampled volumes is $\begin{matrix}{{r( {x,y,z} )} = {{g_{1}( {x,y,z} )}*{g_{2}( {x,y,z} )}}} \\{= {{f( {{x + {\Delta\quad x}},{y + {\Delta\quad y}},{z + {\Delta\quad z}}} )}*{f( {x,y,z} )}*{w_{1}( {x,y,z} )}*}} \\{{w_{2}( {x,y,z} )}*{{\Pi( {x,y,z} )}.}}\end{matrix}$That is just a blurred version of the original correlation; the exactlocation of the maxima is shaped by the blurring function h(x, y,z)=w₁(x, y, z)*w₂(x, y, z). That is, the correlation is distorted by thefunction h(x, y, z). FIG. 4 shows an idealized window function for w₁(x,y, z) and w₂(x, y, z) and its corresponding supporting region of theblurring function, h(x, y, z).

Finding the displacement using the above procedure works fine fornoise-free data, but noise makes the search more difficult. Let g₁(x, y,z)=(f(x, y, z)+n₁(x, y, z))*w₁(x, y, z) and g₂(x, y, z)=(x+Δx, y+Δy,z+Δz)+n₂(x, y, z))*w₂(x, y, z) be the corresponding noisy volumes, wheren₁(x, y, z) and n₂(x, y, z) are two uncorrelated noise sources. Thus,the correlation of g₁ and g₂ is given by $\begin{matrix}{{r( {x,y,z} )} = {{g_{1}( {x,y,z} )}*{g_{2}( {x,y,z} )}}} \\{= \lbrack {{{f( {{x + {\Delta\quad x}},{y + {\Delta\quad y}},{z + {\Delta\quad z}}} )}*{f( {x,y,z} )}} +} } \\{{{f( {{x + {\Delta\quad x}},{y + {\Delta\quad y}},{z + {\Delta\quad z}}} )}*{n_{1}( {x,y,z} )}} + {{f( {x,y,z} )}*}} \\{{ {n_{2}( {x,y,z} )} \rbrack*{h( {x,y,z} )}*{\Pi( {x,y,z} )}},}\end{matrix}$and the maximum is no longer guaranteed to be given by the displacement,especially for functions with smooth autocorrelation functions likestandard MRI. The smooth autocorrelation functions make the error afunction of the noise power. Autocorrelation functions whose shapes arecloser to a Dirac delta function δ(x, y, z) are less sensitive to noise,which is the reason why many registration algorithms work with edges.The autocorrelation function of the gradient magnitude of standard MRIis closer to a Dirac delta function. Therefore, the registration of thegradient is less sensitive to noise.

FIGS. 5A-5C shows a 1D example of the effect of the derivative on theautocorrelation function of a band-limited signal. FIG. 5A shows anoriginal signal f(x). FIG. 5B shows a smooth estimate of the magnitudeof the derivative, namely,g(x)=|f(x)*[−1 −1 0 0 0 1 1]|FIG. 5C shows autocorrelations; the dashed curve represents f*f, whilethe curve shown in crosses represents g*g. For that example, a smoothderivative operator is used to reduce the noise level.

For the above reasons, the automatic registration is based on findingthe maximum on the correlation among the magnitude gradient of the twomagnetic resonance images:$( {{\Delta\quad x},{\Delta\quad y},{\Delta\quad z}} ) = {{Arg}{\max\limits_{({{\Delta\quad x},{\Delta\quad y},{\Delta\quad z}})}{\int{\int{\int{{{\nabla{g_{1}( {\alpha,\beta,\gamma} )}}}{{\nabla{g_{2}( {{{\Delta\quad x} + \alpha},{{\Delta\quad y} + \beta},{{\Delta\quad z} + \gamma}} )}}}{\mathbb{d}\alpha}{\mathbb{d}\beta}{{\mathbb{d}\gamma}.}}}}}}}$

In sampled images the gradient ∇ can be approximated by finitedifferences:${{{\nabla{g( {x,y,z} )}}} = \sqrt{{d_{x}( {x,y,z} )}^{2} + {d_{y}( {x,y,z} )}^{2} + {d_{z}( {x,y,z} )}^{2}}},{where}$${{d_{x}( {x,y,z} )} = \frac{{{l( {x,y,z} )}*{g( {{x + {\delta\quad x}},y,z} )}} - {{l( {x,y,z} )}*{g( {{x - {\delta\quad x}},y,z} )}}}{2\delta\quad x}},{{d_{y}( {x,y,z} )} = \frac{{{l( {x,y,z} )}*{g( {x,{y + {\delta\quad y}},z} )}} - {{l( {x,y,z} )}*{g( {x,{y - {\delta\quad y}},z} )}}}{2\delta\quad y}},{{d_{z}( {x,y,z} )} = \frac{{{l( {x,y,z} )}*{g( {x,y,{z + {\delta\quad z}}} )}} - {{l( {x,y,z} )}*{g( {x,y,{z - {\delta\quad z}}} )}}}{2\delta\quad z}},$where δx, δy, δz are the sampling rates, and l(x, y, z) is a low passfilter used to remove noise from the images and to compensate thedifferences between in-slice sampling and inter-slice sampling.

The maximization can be done using any standard maximization technique.The preferred embodiment uses a simple hill-climbing technique becauseof the small displacements. The hill-climbing technique evaluates thecorrelation at the six orthogonal directions: up, down, left, right,front and back. The direction that has the biggest value is chosen asthe next position. That simple technique works well for the registrationof two orthogonal data sets, as the one expects for involuntary motionduring scans.

To avoid being trapped in local maxima and to speed up the process, amulti-resolution approach can be used. That multi-resolution approachselects the hill-climbing step as half the size of the previous step.Five different resolutions are used. The coarsest resolution selected istwice the in-plane resolution of the system, and the smallest size isjust 25 percent of the in-plane resolution.

Even with a very simple optimization approach, the computation of thecorrelation of the whole data set can be time consuming; therefore, theregistration of two images can take time. To speed up the correlation ofthe two data sets, just a small subsample of the points with very highgradient are selected for use in the correlation process.

Some images suffer from a small rotation. In that case, the algorithm isextended to search for the image rotation. The same hill-climbingtechnique is used to find the rotation between images; but instead ofdoing the search in a three-dimensional space, the algorithm has to lookat a six-dimensional space. That search space includes the threedisplacements and three rotations along each axis. At each step therotation matrix is updated and used to compensate for the small rotationbetween images.

Image Fusion (Step 208)

Once the two images are registered, an isotropic high resolution imageis created from them. Due to the different shape between voxel samplingvolumes (w₂(x, y, z) and w₁(x, y, z)), one has to be careful whenestimating every high resolution voxel value from the input data. Assumethat the first image has been scanned in the x-direction and the secondhas been scanned in the y-direction. Therefore, there is high-resolutioninformation in the z-direction in both images.

FIGS. 6A and 6B show the voxel shapes of the two input images, where thein-slice resolution is equal, and the inter-slice resolution is fourtimes lower. Given that configuration, the problem of filling thehigh-resolution volume is a 2D problem. In a single 4×4-voxel window ofthe high resolution image, as seen in FIG. 6C, then for every 16high-resolution voxels there are only 8 known low-resolution voxels;therefore, that is an ill posed problem.

To address that problem, assume that every high-resolution voxel is justa linear combination of the two low-resolution functions:g(x,y,z)=h ₁(x/s _(d) ,y,z)+h ₂(x,y/s _(d) ,z),where h₁, h₂ are two functions which are back projected in such a waythat${{g_{1}( {{x/s_{d}},y,z} )} = {\sum\limits_{\forall{x \in w_{1}}}{g( {x,y,z} )}}},{{g_{2}( {x,{y/s_{d}},z} )} = {\sum\limits_{\forall{y \in w_{2}}}{{g( {x,y,z} )}.}}}$That represents a linear system with the same number of knowns asunknowns. The known values are the observed image voxels, while thevalues to back-project which match the observation are estimated. Noiseand inhomogeneous sampling make the problem a little bit harder; butthat linear system can efficiently be solved using projection on convexsets (POCS). Although, in theory, all the components have to beorthogonally projected, it can be shown that the following projectingscheme also works:${{h_{1}^{k + 1}( {{x/s_{d}},y,z} )} = {{h_{1}^{k}( {{x/s_{d}},y,z} )} - {\alpha\frac{( {{{nh}_{1}^{k}( {{x/s_{d}},y,z} )} + {\sum\limits_{\forall{y \in w_{1}}}{h_{2}^{k}( {x,,{y/s_{d}},z} )}} - {g_{1}( {{x/s_{d}},y,z} )}} )n}{n^{2} + m}}}},{{h_{2}^{k + 1}( {x,{y/s_{d}},z} )} = {{h_{2}^{k}( {x,{y/s_{d}},z} )} - {\alpha\frac{( {{{mh}_{2}^{k}( {x,{y/s_{d}},z} )} + {\sum\limits_{\forall{y \in w_{2}}}{h_{1}^{k}( {{x/s_{d}},y,z} )}} - {g_{2}( {x,{y/s_{d}},z} )}} )m}{n + m^{2}}}}},$where 0<α<1, n=the window size of w₁, m=the window size of w₂, h₁ ⁰(x,y, z)=g(x, y, z) and h₂ ⁰(x, y, z)=g₂(x, y, z) are the initial guessesfor the estimation of back-projected functions. The advantage of thatapproach over standard orthogonal projection is that is equations aresimpler and that they can be implemented efficiently on a computer.

Experimental Data

Some experimental data produced by the above technique will now bedescribed.

FIGS. 7A, 7B and 7C respectively show the original MRI sagittal scan,the original MRI axial scan and the fused image of a human shoulder seenalong an axial view. FIGS. 7D, 7E and 7F show the same, except seenalong a sagittal view. FIGS. 7G, 7H and 7I show the same, except seenalong a coronal view. In all three cases, the image is noticeablyimproved.

FIGS. 8A, 8B and 8C show axial, sagittal and coronal slices,respectively, of simple fusion without registration. FIGS. 8D, 8E and 8Fshow the same slices with simple fusion after registration. FIGS. 8G, 8Hand 8I show the same slices with complete image fusion. The simplefusion is g(x, y, z)=0.5 g₁(x, y, z)+0.5 g₂(x, y, z), and the completefusion is g(x, y, z)=h₁(x, y, z)+h₂(x, y, z), wherein h₁ and h₂ are thetwo functions which minimize the reconstruction error.

FIGS. 9A-9I show fusion of orthogonal images without correlation. FIGS.9A, 9B and 9C show axial views of the original MRI sagittal scan, theoriginal axial scan and the fused image, respectively, for an axialview. FIGS. 9D, 9E and 9F show the same for a sagittal view. FIGS. 9G,9H and 9I show the same for a coronal view.

Multiple Local Coil Receivers and Multispectral Imaging

Good signal-to-noise ratio is very important for an unsupervisedsegmentation algorithm. Even more important is the contrast-to-noiseratio among neighboring tissues. When local receiving coils are used,the signal from points far from the coil location is weak; therefore,contrast among tissues located far from the receiving coil is low. Someresearchers have proposed several software alternatives to correct thissignal fading, but this will increase the noise levels as well. Thus, itwill not solve the problem. The preferred embodiment uses two or morereceiving coils, which will improve the signal reception at farlocations.

FIGS. 10A and 10B show the advantage of using multiple coils. FIG. 10Ashows an MRI image of a knee using three surface coils. FIG. 10B showsan MRI image of the same knee using four surface coils.

Multispectral images will now be considered. Such images can beanalogized to multispectral optical images, in which, for example, red,blue and green images are combined to create a single color image.

FIGS. 11A and 11B show a two-band spectral image of a knee. FIG. 11Ashows a cross section of a fat suppression MRI scan of the knee, wherefat, and bone tissues have almost the same low density, cartilage has avery high density, and muscle tissue has a medium density. FIG. 11Bshows the same knee, but now, muscle tissue and cartilage have the samedensity, making them very hard to differentiate. Those images clearlyshow the advantage of the multispectral image approach in describing theanatomy.

The analysis of a multispectral image is more complex than that of asingle-spectrum image. One way to simplify the analysis is to reduce thenumber of bands by transforming an N-band multispectral image in such away that passes the most relevant image into an (N-n)-band image. Thetransform that minimizes the square error between the (N-n)-band imageand the N-band image is the discrete Karhuen-Loeve (K-L) transform. Theresulting individual images from the transformed spectral image afterapplying the K-L transform are typically known as the principalcomponents of the image.

Let the voxel x be an N-dimensional vector whose elements are the voxeldensity from each individual pulse sequence. Then the vector mean valueof the image is defined as${m_{x} = {{E\lbrack x\rbrack} = {\frac{1}{M}{\sum\limits_{k = 1}^{M}x_{k}}}}},$and the covariance matrix is defined as${C_{x} = {{E\lbrack {( {x - m_{x}} )( {x - m_{x}} )^{2}} \rbrack} = {{\frac{1}{M}{\sum\limits_{k = 1}^{M}{x_{k}x_{k}^{T}}}} - {m_{k}m_{k}^{T}}}}},$where M is the number of voxels in the image.

Because C_(x) is real and symmetric, a set of n orthogonal eigenvectorscan be found. Let e_(i) and λ_(i), i=1, 2, . . . N, be the eigenvectorsand the corresponding eigenvalues of C_(x), so that λ_(j)>λ_(j)+1. Let Abe the matrix formed with the eigenvectors of C_(x). Then thetransformationy=A(x−m)_(x)is the discrete K-L transform, and y_(i), i=1, 2, . . . N, are thecomponents of a multispectral image. FIGS. 11C and 11D show theprincipal components of the two-band spectral image shown in FIGS. 11Aand 11B. The advantage of the K-L decomposition of a multispectral imageis that the image with the highest contrast is associated with thehighest eigenvalue of the correlation matrix, and the image associatedwith the smallest eigenvalues usually is irrelevant.

The high-contrast, multispectral images thus formed can be utilized fordiagnosis and for input to post-processing systems, such asthree-dimensional rendering and visualization systems. If themultispectral data are from orthogonal planes or are acquired with somemisregistration, the registration and orthogonal fusion steps describedherein can be employed to enhance the resolution and contrast.

While a preferred embodiment of the present invention has been set forthabove, those skilled in the art who have reviewed the present disclosurewill readily appreciate that other embodiments can be realized withinthe scope of the present invention. For example, while the invention hasbeen disclosed as used with the hardware of FIG. 1, other suitable MRIhardware can be used. For that matter, the invention can be adapted toimaging techniques other than MRI, such as tomography. Also, while thescans are disclosed as being in orthogonal directions, they can be takenin two different but non-orthogonal directions. Therefore, the presentinvention should be construed as limited only by the appended claims.

1-20. (canceled)
 21. A method of forming an image of a subject, themethod comprising: (a) performing an MRI scan on the subject to takeimage data having a plurality of spectral bands; and (b) forming theimage from the image data.
 22. The method of claim 21, wherein step (b)comprises selecting a subplurality of the plurality of spectral bands toform the image.
 23. The method of claim 22, wherein the subplurality ofthe plurality of spectral bands is selected by ranking the plurality ofspectral bands in order of image contrast and selecting the spectralbands whose image contrast is highest.
 24. The method of claim 23,wherein the plurality of spectral bands is ranked in order of imagecontrast by: deriving a covariance matrix from the plurality of spectralbands; deriving a set of orthogonal eigenvectors and a corresponding setof eigenvalues from the covariance matrix; and ranking the orthogonaleigenvectors in order of their corresponding eigenvalues.
 25. The methodof claim 21, wherein step (a) is performed with a plurality of receivingcoils.
 26. The method of claim 25, wherein step (a) is performed usingat least three receiving coils.
 27. The method of claim 26, wherein step(a) is performed using at least four receiving coils. 28-47. (canceled)48. A system for forming an image of a subject, the system comprising:scanning means for performing an MRI scan on the subject to take imagedata having a plurality of spectral bands; and computing means forforming the image from the image data.
 49. The system of claim 48,wherein the computing means selects a subplurality of the plurality ofspectral bands to form the image.
 50. The system of claim 49, whereinthe subplurality of the plurality of spectral bands is selected byranking the plurality of spectral bands in order of image contrast andselecting the spectral bands whose image contrast is highest.
 51. Thesystem of claim 50, wherein the plurality of spectral bands is ranked inorder of image contrast by: deriving a covariance matrix from theplurality of spectral bands; deriving a set of orthogonal eigenvectorsand a corresponding set of eigenvalues from the covariance matrix; andranking the orthogonal eigenvectors in order of their correspondingeigenvalues.
 52. The system of claim 48, wherein the scanning meanscomprises a plurality of receiving coils.
 53. The system of claim 52,wherein the plurality of receiving coils comprises at least threereceiving coils.
 54. The system of claim 53, wherein the plurality ofreceiving coils comprises at least four receiving coils.